Stent fabricated from polymer composite toughened by a dispersed phase

ABSTRACT

Stents fabricated from polymer composites toughened by a dispersed phase are disclosed.

CROSS REFERENCE TO RELATED APPLICATION

This application is a divisional of U.S. patent application Ser. No.11/827,180 filed Jul. 10, 2007 which claims the benefit of U.S. PatentApplication No. 60/830,211 which was filed on Jul. 11, 2006, both ofwhich are incorporated by reference herein.

BACKGROUND OF THE INVENTION

Field of the Invention

This invention relates to a stent fabricated at least in part from apolymer composite toughened by a dispersed polymer phase.

Description of the State of the Art

This invention relates to radially expandable endoprostheses, which areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices, which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty in the vascular system. “Restenosis”refers to the reoccurrence of stenosis in a blood vessel or heart valveafter it has been treated (as by balloon angioplasty, stenting, orvalvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment. “Deployment” corresponds tothe expanding of the stent within the lumen at the treatment region.Delivery and deployment of a stent are accomplished by positioning thestent about one end of a catheter, inserting the end of the catheterthrough the skin into a bodily lumen, advancing the catheter in thebodily lumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a constraining member such as aretractable sheath or a sock. When the stent is in a desired bodilylocation, the sheath may be withdrawn which allows the stent toself-expand.

The stent must be able to satisfy a number of mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. Therefore, a stent must possess adequate radialstrength. Radial strength, which is the ability of a stent to resistradial compressive forces, is due to strength and rigidity around acircumferential direction of the stent. Radial strength and rigidity,therefore, may also be described as, hoop or circumferential strengthand rigidity.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.This service life for a biodegradable stent is the length of time neededto support the vessel to prevent vessel recoil and negative remodeling.For example, a radially directed force may tend to cause a stent torecoil inward. Generally, it is desirable to minimize recoil. Inaddition, the stent must possess sufficient flexibility to allow forcrimping, expansion, and cyclic loading. Longitudinal flexibility isimportant to allow the stent to be maneuvered through a tortuousvascular path and to enable it to conform to a deployment site that maynot be linear or may be subject to flexure. Finally, the stent must bebiocompatible so as not to trigger any adverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. In other embodiments, the scaffolding can be formedfrom machining, or cutting a pattern out of tubing. The scaffolding isdesigned so that the stent can be radially compressed (to allowcrimping) and radially expanded (to allow deployment). A conventionalstent is allowed to expand and contract through movement of individualstructural elements of a pattern with respect to each other.

Additionally, a medicated stent may be fabricated by coating the surfaceof either a metallic or polymeric scaffolding with a polymeric carrierthat includes an active or bioactive agent or drug. Polymericscaffolding may also serve as a carrier of an active agent or drug.

Furthermore, it may be desirable for a stent to be biodegradable. Inmany treatment applications, the presence of a stent in a body may benecessary for a limited period of time until its intended function of,for example, maintaining vascular patency and/or drug delivery isaccomplished. Therefore, stents fabricated from biodegradable,bioabsorbable, and/or bioerodable materials such as bioabsorbablepolymers should be configured to completely erode only after theclinical need for them has ended.

Potential problems with biodegradable polymeric implantable medicaldevices, such as stents, include insufficient toughness and slowdegradation rate.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a stent comprisinga body fabricated from a bioabsorbable polymer composite, the polymercomposite comprising: a high toughness polymer dispersed within a matrixpolymer, the matrix polymer being glassy at physiological conditions,wherein the high toughness polymer enhances the fracture toughness ofthe composite at physiological conditions.

Further embodiments of the present invention include a stent comprisinga composite layer formed from a bioabsorbable polymer composite, thepolymer composite comprising: a high toughness polymer dispersed withina matrix polymer, the matrix polymer being glassy at physiologicalconditions, wherein the high toughness polymer enhances the fracturetoughness of the composite at physiological conditions.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A depicts a view of a stent.

FIG. 1B depicts a section of a structural element from the stentdepicted in FIG. 1A.

FIG. 2 depicts a schematic close-up view of the section depicted in FIG.1B.

FIG. 3 depicts an axial cross-section of a strut showing a coating overa scaffolding or body.

FIG. 4 depicts an exemplary axial cross-section of a strut with anabluminal composite layer over a luminal stent body layer.

FIG. 5 depicts an exemplary axial cross-section of a strut with a stentbody layer between a luminal composite layer and an abluminal compositelayer.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of the present invention include a stent fabricatedat least in part from a polymer-polymer composite that includes adispersed polymer phase. The dispersed phase tends to enhance thetoughness of the composite.

Embodiments of the present invention relate to implantable medicaldevices including, but is not limited to, self-expandable stents,balloon-expandable stents, stent-grafts, other expandable tubulardevices for various bodily lumen or orifices. Such devices can bedesigned for the localized delivery of a therapeutic agent. A medicatedimplantable medical device may be constructed by coating the device orsubstrate with a coating material containing a therapeutic agent. Thesubstrate of the device may also contain a therapeutic agent.

FIG. 1A depicts a view of a stent 100. In some embodiments, a stent mayinclude a pattern or network of interconnecting structural elements 105.Stent 100 may be formed from a tube (not shown). The pattern ofstructural elements 105 can take on a variety of patterns. Thestructural pattern of the device can be of virtually any design. Theembodiments disclosed herein are not limited to stents or to the stentpattern illustrated in FIG. 1A. The embodiments are easily applicable toother patterns and other devices. The variations in the structure ofpatterns are virtually unlimited. A stent such as stent 100 may befabricated from a tube by forming a pattern with a technique such aslaser cutting or chemical etching.

An implantable medical device can be made partially or completely from abiodegradable, bioabsorbable, biostable polymer, or a combinationthereof. A polymer for use in fabricating an implantable medical devicecan be biostable, bioabsorbable, biodegradable or bioerodable. Biostablerefers to polymers that are not biodegradable. The terms biodegradable,bioabsorbable, and bioerodable are used interchangeably and refer topolymers that are capable of being completely degraded and/or erodedwhen exposed to bodily fluids such as blood and can be graduallyresorbed, absorbed, and/or eliminated by the body. The processes ofbreaking down and absorption of the polymer can be caused by, forexample, hydrolysis and metabolic processes.

Stents are typically subjected to stress during use. “Use” includesmanufacturing, assembling (e.g., crimping a stent on balloon), deliveryof a stent through a bodily lumen to a treatment site, deployment of astent at a treatment site, and treatment after deployment. Both theunderlying scaffolding or substrate and the coating experience stressthat result in strain in the substrate and coating. In particular,localized portions of the stent's structure undergo substantialdeformation, such as at the apex regions of bending elements andexperience relatively high stress and strain during crimping, expansion,and after expansion of the stent.

As discussed above, it is important for a stent body or scaffolding tohave high radial strength and stiffness so that it can a support alumen. Some crystalline or semi-crystalline polymers that are glassy orhave a Tg above body temperature are particularly attractive as stentmaterials due to their strength and stiffness. Some of these polymersthat may be suitable for implantable medical devices such as stents havepotential shortcomings. One shortcoming of such polymers is that theirtoughness can be lower than desired, in particular, for use in stentapplications. For example, polymers such as poly(L-lactide) (PLLA) arestiff and strong, but tend to be brittle under physiological conditions.Physiological conditions refer to conditions that an implant is exposedto within a human body. Physiological conditions include, but arelimited to, human body temperature, approximately 37° C. These polymerscan exhibit a brittle fracture mechanism at these conditions in whichthere is little or no plastic deformation prior to failure. As a result,a stent fabricated from such polymers can have insufficient toughnessfor the range of use of a stent.

As discussed above, a medicated implantable medical device, such as astent, may be fabricated by coating the surface of a stent with a drug.For example, a device can have a coating including a drug dispersed in apolymeric carrier disposed over a substrate of the stent. Suchpolymer-based coatings may be particularly vulnerable to mechanicalinstability during use of a stent. Such mechanical instability forcoatings can include fracture and detachment from a substrate, forexampling, peeling. Some polymers may be susceptible to such mechanicalinstability due to insufficient toughness at high deformations. Thus, itis important for a polymer-based coating to be tough and have a highresistance to cracking in the range of deformations that occur duringcrimping, during deployment of a stent, and after deployment.

Furthermore, some biodegradable polymers have a degradation rate that isslower than desired for certain stent treatments. As a result, thedegradation time of a stent made from such polymers can be longer thandesired. For example, a stent made from a semicrystalline polymer suchas PLLA can have a degradation time between about two and three years.In some treatment situations, a shorter degradation time is desirable,for example, less than a year.

One way to form a tougher polymeric material from a brittle polymer isby making a composite including the brittle polymer and another polymerthat has a higher fracture toughness than the brittle polymer. Thehigher toughness polymer should also be immiscible with or form aseparate phase from the brittle polymer. For example, the highertoughness polymer can be dispersed as discrete phase domains within thematrix polymer. The fracture toughness of the composite is increasedsince the dispersed phase can absorb energy arising from stress impartedto a part made from the composite. The increase in the fracturetoughness can be enhanced by increasing the adhesion of the dispersedphase with the continuous polymer phase. To ensure good energy transferbetween interfaces of the phases, it is important that there besufficient bonding or adhesion between the phases. See, Y. Wang, etc.Journal of Polymer Science Part A: Polymer Chemistry, 39, 2001,2755-2766.

Certain embodiments of the present invention include a stent having abody or scaffolding fabricated from a bioabsorbable polymer composite.In some embodiments, the polymer composite includes a high toughnesspolymer dispersed within a matrix polymer. The high toughness polymercan be a dispersed phase within the matrix polymer which can be acontinuous polymer phase. The matrix polymer may be glassy atphysiological conditions. In an embodiment, the high toughness polymermay have a lower modulus than the glassy matrix polymer.

In some embodiments, a stent body can refer to a stent scaffolding withan outer surface to which no coating or layer of material different fromthat of which the device is manufactured has yet been applied. If thebody is manufactured by a coating process, the stent body can refer to astate prior to application of optional additional coating layers ofdifferent material. By “outer surface” is meant any surface howeverspatially oriented that is in contact with bodily tissue or fluids. Astent body can refer to a stent scaffolding formed by laser cutting apattern into a tube or a sheet that has been rolled into a cylindricalshape.

In some embodiments, a majority, substantially all, or all of the stentbody or scaffolding can be made from the composite. Substantially all ofthe body can refer to greater than 90%, 95%, or greater than 99% of thebody.

Additionally, the matrix polymer is stronger and stiffer than the hightoughness polymer and is primarily or completely responsible forproviding strength required to support the walls of a bodily lumen whenthe stent is deployed at a treatment site. In such embodiments, the hightoughness polymer enhances the fracture toughness of the composite atphysiological conditions. Thus, the high toughness polymer reduces orprevents formation of cracks during use of the stent. In someembodiments, the interfacial adhesion of the dispersed phase with thematrix or continuous polymer phase is high enough to allow for thedispersed phase to increase the fracture toughness of the compositeduring use of a stent. “Use” includes manufacturing, assembling (e.g.,crimping a stent on balloon), delivery of a stent through a bodily lumento a treatment site, and deployment of a stent at a treatment site. Insuch embodiments, the interfacial adhesion between the dispersed phaseof the high toughness polymer and the matrix polymer is high enough suchthat when a stress is placed upon the interface during use of a stent,the high toughness polymer fails before an interfacial bond between thephases. In other such embodiments, the interfacial adhesion between thedispersed phase of the high toughness polymer and the matrix polymer isgreater than the strength of the high toughness polymer. In additionalembodiments, the interfacial adhesion between the dispersed phase of thehigh toughness polymer and the matrix polymer is at least 10% of thestrength of the high toughness polymer.

FIG. 1B depicts a section of a segment 110 of strut 105 from the stentdepicted in FIG. 1A. FIG. 2 depicts a microscopic schematic view of aportion 140 of segment 110 of a strut as depicted in FIG. 1B. Portion140 includes a dispersed phase with a plurality of discrete polymerphase regions 200 dispersed within a continuous polymer phase 210.Discrete phase regions include a polymer that is tougher than thepolymer of the continuous polymer phase. L_(D) is a characteristicdimension of discrete phase regions 200.

In further embodiments of the present invention, a stent can includelayers of bioabsorbable composite with a high toughness polymerdispersed within a matrix polymer. In such embodiments, a compositelayer can be therapeutic with a drug or active agent incorporated withinthe layer. In some embodiments, the composite layer can be formed aboveor over at least a portion of a stent body. The stent body can bepolymeric, metallic, or a combination thereof. The high toughnesspolymer enhances the fracture toughness of the composite atphysiological conditions which reduces or prevents fracture anddetachment of the composite layer from a stent.

In some embodiments, the composite layer is a coating layer that may beformed above the stent body or scaffolding. The stent body is primarilyor completely responsible for mechanical support of a bodily lumen whenthe stent is deployed. In an embodiment, the stent body or scaffoldingcan be composed of a polymer that allows the stent body to providerequisite mechanical support to a bodily lumen. In an embodiment, astent body or scaffolding is composed of a strong, glassy polymer.

In some embodiments, the stent body or scaffolding is also composed of acomposite, as described above. In this embodiment, the stent bodycomposite can be stronger and stiffer than the composite coating layer.The coating layer can have a higher weight percent of high toughnesspolymer than the stent body composite. In certain embodiments, the hightoughness polymer can be all or a majority of the composite of a layerover a stent body.

In certain embodiments, a composite layer may be above all or a portionof a stent body or scaffolding. FIG. 3 depicts an axial cross-section ofa strut 300 showing a coating 305 over a stent scaffolding or body 310.Coating 305 is above a luminal surface 315, abluminal surface 320, andsidewall surfaces 325 of body 310. In another embodiment, a compositelayer can be topcoat layer disposed over polymer and drug coating layer.The topcoat layer can be used to control the drug release from thepolymer and drug layer.

In further embodiments, a stent can include structural elements havingan abluminal luminal, or both abluminal and luminal composite layers. Inone embodiment, a stent has an abluminal composite layer over a stentbody polymer layer capable of providing mechanical support of a bodilylumen. In another embodiment, a stent has a luminal composite layer andan abluminal glassy polymer layer. In an additional embodiment, thestent has luminal and abluminal composite layers with a stent body layerbetween the abluminal and luminal composite layers. The stent body layercan be glassy polymer or a composite layer that can be different fromthe abluminal/luminal composite layer. In such embodiments, the stentbody polymer layer is primarily or completely responsible for providingmechanical support to a bodily lumen when the stent is deployed. FIG. 4depicts an exemplary axial cross-section of a strut 400 with a compositelayer 405 over a luminal surface 415 of a stent body layer 410. FIG. 5depicts an exemplary axial cross-section of a strut 500 with a stentbody layer 510 between a luminal composite layer 505 and an abluminalcomposite layer 515.

In some embodiments, the matrix polymer of the composite is acrystalline or semicrystalline polymer having a degree of crystallinitygreater than about 30%. Exemplary matrix polymers include PLLA,polyglycolide (PGA), poly(L-lactide-co-D,L-lactide),poly(L-lactide-co-glycolide), poly(L-lactide-co-caprolactone),poly(L-lactide-co-trimethylene carbonate), poly(ester amide) (PEA), orcopolymers thereof.

In some embodiments, the high toughness polymer exhibits a rubbery orelastomeric behavior at physiological conditions. An “elastomeric” or“rubbery” polymer refers to a polymer that exhibits elastic deformationthrough all or most of a range of deformation. Such elastomericproperties provide the composite with a high fracture toughness duringuse of the stent. In some embodiments, the high toughness polymer hasglass transition temperature (Tg) below body temperature. Additionally,the high toughness polymer may be completely or substantially amorphous.

In some embodiments, the high toughness polymer is rubbery,bioabsorbable or biodegradable, biocompatible polymer. Exemplarybiodegradable polymers that are elastomeric or rubbery at physiologicalconditions include, but are not limited to, poly(butylene succinate)(PBS), polycaprolactone (PCL), poly(trimethylene carbonate) (PTMC),poly(4-hydroxy butyrate) (PHB), aliphatic polyanhydrides,polyorthoesters, and polydioxanone (PDO). An exemplary commercialembodiment of PBS is Bionolle® 3000 from Showa Highpolymer Co. Ltd.,Tokyo, Japan. In other embodiments, the high toughness polymer can be arandom or alternating copolymer of the above polymer or other rubberypolymers.

The weight percent of the high toughness polymer in the composite can beadjusted to obtain a desired or optimal degree of fracture toughness. Inparticular, a weight percent of high toughness polymer can be selectedto result in a selected maximum threshold number of cracks or no cracksin a stent body or coating layer upon crimping, deployment, or both. Inaddition, a weight percent of high toughness polymer can be selected toresult in no or substantially no detachment of a coating layer uponcrimping, deployment, or both. In addition, the amount of the hightoughness polymer can be limited by desired or required radial strengthof the stent. In some embodiments, the high toughness polymer can be atleast 0.1, 1, 3, 5, 10, 20, 30, or 40 wt % of the composite.

An exemplary embodiment of a stent body or scaffolding can be fabricatedfrom a composite including a blend of PBS dispersed within PLLA. Theweight percent of PBS can be at least 20, 40, 60, 80 wt % or greaterthan 80 wt %. An exemplary embodiment of a coating for a stent body is ablend of PBS dispersed within PDLA with a weight percent of PBS of atleast 20, 40, 60, 80 wt % or greater than 80 wt %.

In additional embodiments, the high toughness polymer is a random oralternating copolymer including elastomeric units and glassy polymerunits. In other embodiments, the high toughness polymer can be a blockcopolymer including elastomeric blocks and glassy polymer blocks. Therelative weight percent of the elastomeric and glassy polymer units orblocks can be adjusted to obtain desired properties of the polymer, suchas, modulus, Tg, and elastomeric behavior. It is expected that as theweight percent of elastomeric units or blocks increase in the polymer,the modulus of the high toughness polymer decreases and the flexibilityincreases. Additionally, the weight percent of the elastomeric units orblocks can be adjusted so that the high toughness polymer is elastomericat physiological conditions and/or the Tg of the high toughness polymeris below body temperature.

Exemplary elastomeric units include butylene succinate (BS),caprolactone (CL), trimethyl carbonate (TMC), 4-hydroxy butyrate (HB),and dioxanone (DO). Exemplary glassy polymer units include L-lactide(LLA), and glycolide (GA). Exemplary block copolymers that can be usedas a high toughness polymer can include blocks of these elastomericunits and blocks of these glassy units.

In some embodiments, a random, alternating, or block high toughnesscopolymer can have glassy units or blocks that are the same as thematrix polymer. For example, the matrix polymer can be PLLA and the hightoughness polymer can be P(LLA-co-TMC). In such embodiments, theadhesion of the dispersed phase to the matrix polymer can be enhanceddue to the compatibility of the glassy polymer with the matrix polymer.In some embodiments, blocks or segments of the high toughness polymercan phase separate into the matrix polymer. It is believed that suchphase separation increases the interfacial adhesion of the dispersedphase with the continuous phase.

Additionally, the weight percent of the glassy and elastomeric units orblocks of the high toughness polymer can be adjusted so that the hightoughness polymer is immiscible with the matrix polymer. In someembodiments, the elastomeric units can be at least 0.1, 1, 3, 5, 10, 20,30, 40 wt % or more than 40 wt % of the high toughness copolymer.

In an exemplary embodiment, a stent body can be composed of a PLLAmatrix with dispersed 70/30 P(LLA-co-TMC) (70 wt % LLA and 30 wt % TMC).An exemplary coating or outer layer of a stent can be apoly(D,L-lactide) (PDLA) matrix with dispersed 70/30 P(DLA-co-TMC) (70wt % DLA and 30 wt % TMC).

In some embodiments, the adhesion of the dispersed phase with the matrixor continuous polymer phase can be enhanced by including acompatibilizer in the composite. In general, a “compatibilizer” refersto an interfacial agent that modifies the properties of an immisciblepolymer blend or composite which facilitates formation of uniform blend,and increases interfacial adhesion between the phases. Compatibilizationrefers to the process of modification of the interfacial properties inan immiscible polymer blend that results in formation of interphases(region of concentration gradient between phases) and stabilization ofthe morphology. In some embodiments, a compatibilizer can be a blockcopolymer including blocks that are miscible with the matrix polymer andblocks that are miscible with the high toughness polymer of thedispersed phase. In one such embodiment, the compatibilizer can includeglassy blocks of the matrix polymer and elastomeric blocks of the hightoughness polymer. In an exemplary embodiment, the matrix polymer can bePLLA, the high toughness polymer can be PTMC and the compatibilizer canbe PLLA-b-PTMC.

In some embodiments, the high toughness polymer can have a fasterdegradation rate than the matrix polymer. As indicated above, a polymer,such as PLLA, can have a degradation rate that is slower than desiredfor certain stent treatments. The slow degradation rate is due at leastin part to the crystallinity of a matrix polymer. The faster degradationrate of the high toughness polymer can be due at least in part to alower degree of crystallinity or a higher percentage amorphous structureof the dispersed phase. The diffusion rate of fluids through anamorphous structure is generally faster than through a crystallinestructure. The faster degrading, high toughness polymer increases waterpenetration and content in the dispersed phase and matrix polymer phase.The increased water penetration and content causes an increase in thedegradation rate of the composite. As a result, the degradation time ofa stent body or coating is decreased.

It is believed that when a device is placed under stress, the dispersedphase tends to absorb energy when a fracture starts to propagate througha structural element. Thus, crack propagation through the continuousphase may then be reduced, inhibited, or eliminated. As a result,fracture toughness of the polymer composite, and thus, the stent body orcoating tends to be increased.

Generally, it is desirable for the dispersed phase to be uniformly orsubstantially uniformly dispersed throughout the polymer matrix tofacilitate the increase in toughness. The more dispersed the dispersedphase, the greater is the increase in toughness. Additionally, theincrease in toughness is related to the size of the discrete phaseregions. The characteristic length of a discrete phase can be 1 nm to100 nm, 100 nm to 500 nm, 500 nm to 1,000 nm, or greater than 1,000 nm.

In certain embodiments, the composite of a high toughness polymer and amatrix polymer can be formed by solution blending, melt blending, or acombination thereof. In some embodiments, a high toughness polymer andmatrix polymer can be melt blended in a mixing apparatus such as anextruder. Representative examples of extruders include, but are notlimited to, single screw extruders, intermeshing co-rotating andcounter-rotating twin-screw extruders, and other multiple screwmasticating extruders. The mixing in the extruder can be sufficient todisperse the high toughness polymer uniformly or relatively uniformlywithin the matrix polymer. Additionally, the mixing can also reduce thesize of the discrete phase domains.

As indicated above, a stent body can be formed from a tube or a sheet.In some embodiments, a tube or sheet can be formed from the compositeusing an extruder or injection molding. The sheet can be rolled orbonded to form a tube. A stent body or scaffolding can then be formedfrom the composite tube by laser machining a stent pattern in the tube.In

In further embodiments, a composite layer over a stent body orscaffolding can be formed by coating the stent body or scaffolding. Insuch embodiments, a coating layer may be formed by applying a coatingmaterial to a body of a stent. The coating material can be a polymersolution that includes the matrix polymer and high toughness polymerdissolved in solvent. The solution can further include a drug dispersedin the solution. The coating material may be applied to the stent bodyby immersing the stent in the coating material, by spraying the materialonto the stent, or by other methods known in the art. The solvent in thesolution is then removed, for example, by evaporation, leaving on thestent surfaces a polymer coating impregnated with the drug.

In another embodiment, a stent having a glassy polymer layer withabluminal, luminal, or both luminal and abluminal composite layers canbe formed from a tube with composite layers. Such a tube can be formedby co-extruding the composite polymer with a polymer that will form alayer that provides mechanical support. A stent can be cut from the tubeto form a layered stent. An active agent or drug can be included in thecomposite layers during extrusion. In some embodiments, a polymersolvent solution can be extruded at a temperature less than atemperature at which the active agent or drug degrades. For example, thecoextrusion can be performed at a temperature below 80° C. or 100° C.

In alternative embodiments, an abluminal or luminal layer can byselectively coated on an abluminal or luminal surface of a stent body.In one embodiment, a controlled deposition system can be used thatapplies various substances only to certain targeted portions of a stent.A representative example of such a system, and a method of using thesame, is described in U.S. Pat. No. 6,395,326 to Castro et al.Alternatively, a luminal or abluminal surface can be masked during thecoating process to selectively coat an abluminal or luminal surface,respectively.

In general, representative examples of polymers that may be used inembodiments of the present invention include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(3-hydroxyvalerate),poly(D,L-lactide-co-glycolide), poly(3-hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(L-lactide-co-glycolide);poly(D,L-lactide), poly(L-lactide-co-D,L-lactide),poly(L-lactide-co-caprolactone), poly(caprolactone), poly(trimethylenecarbonate), polyethylene amide, polyethylene acrylate, poly(glycolicacid-co-trimethylene carbonate), poly(glycolide-co-trimethylenecarbonate), co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes,biomolecules (such as fibrin, fibrinogen, cellulose, starch, collagenand hyaluronic acid), polyurethanes, silicones, polyesters, polyolefins,polyisobutylene and ethylene-alphaolefin copolymers, acrylic polymersand copolymers other than polyacrylates, vinyl halide polymers andcopolymers (such as polyvinyl chloride), polyvinyl ethers (such aspolyvinyl methyl ether), polyvinylidene halides (such as polyvinylidenechloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics(such as polystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose.

Additional representative examples of polymers that may be especiallywell suited for use in embodiments of the present invention includeethylene vinyl alcohol copolymer (commonly known by the generic nameEVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluoropropylene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

For the purposes of the present invention, the following terms anddefinitions apply:

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. True stress denotes the stress where force and area aremeasured at the same time. Conventional stress, as applied to tensionand compression tests, is force divided by the original gauge length.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that results from the applied force. For example, amaterial has both a tensile and a compressive modulus. A material with arelatively high modulus tends to be stiff or rigid. Conversely, amaterial with a relatively high toughness tends to be flexible. Themodulus of a material depends on the molecular composition andstructure, temperature of the material, amount of deformation, and thestrain rate or rate of deformation. For example, below its Tg, a polymertends to be brittle with a high modulus. As the temperature of a polymeris increased from below to above its Tg, its modulus decreases.

“Strain” refers to the amount of elongation or compression that occursin a material at a given stress or load.

“Elongation” may be defined as the increase in length in a materialwhich occurs when subjected to stress. It is typically expressed as apercentage of the original length.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. Thus, a brittle material tends to havea relatively low toughness.

“Solvent” is defined as a substance capable of dissolving or dispersingone or more other substances or capable of at least partially dissolvingor dispersing the substance(s) to form a uniformly dispersed solution atthe molecular- or ionic-size level at a selected temperature andpressure. The solvent should be capable of dissolving at least 0.1 mg ofthe polymer in 1 ml of the solvent, and more narrowly at least 0.5 mg in1 ml at the selected temperature and pressure, for example, ambienttemperature and ambient pressure.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A balloon expandable stent mounted about aballoon disposed on a catheter, the stent comprising: a scaffoldincluding a pattern of interconnecting structural elements, wherein thescaffold is fabricated from a blend of poly(L-lactide) (PLLA) matrixpolymer, a random copolymer of poly(L-lactide-co-caprolactone), and acompatibilizer which is a block copolymer of poly(L-lactide) andpolycaprolactone, and wherein the poly(L-lactide-co-caprolactone) randomcopolymer is at least 10 wt % of the blend and the caprolactone is atleast 20 wt % of the poly(L-lactide-co-caprolactone) random copolymerand; a coating on the scaffold, wherein the coating comprises a drugdispersed within a polymeric carrier.
 2. The stent of claim 1, whereinthe poly(L-lactide-co-caprolactone) forms a dispersed phase within thePLLA matrix polymer.
 3. The stent of claim 2, wherein a characteristiclength of discrete phase regions of the dispersed phase is 100 nm to1000 nm.
 4. The stent of claim 1, wherein greater than 95% of thescaffold is made of the blend.